1. Field of the Invention
The present invention in general concerns magnetic resonance tomography (MRT) as applied in medicine to examine patients. The present invention concerns in particular concerns a method to determine the ADC coefficients in diffusion-weighted magnetic resonance imaging.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear magnetic resonance and has been successfully used as an imaging method for over 15 years in medicine and biophysics. In this examination modality, the subject is exposed to a strong, constant magnetic field. The nuclear spins of the atoms in the subject, which were previously randomly oriented, thereby align. Radio-frequency energy can now excite these “ordered” nuclear spins to a specific oscillation. This oscillation generates the actual measurement signal that is acquired by appropriate reception coils. By the use of inhomogeneous magnetic fields generated by gradient coils, the signals from the measurement subject can be spatially coded in all three spatial directions, what is generally known as “spatial coding”.
In the assessment of pathophysiological events, in particular in the human brain (for example given a stroke), a relatively new MR technique has proven to be particularly effective: diffusion-weighted magnetic resonance tomography.
Diffusion ensues by the thermal translation movement of molecules. It is a random process that is also known as Brownian molecular motion. The distances traveled by the molecules observed in diffusion-weighted MRT measurements are very small; for example, unrestricted diffused water molecules typically in any arbitrary direction over a distance of approximately 20 μm in 100 ms or 60 μm in 1 s. These distances lie in the range of individual cells, in particular in human tissue. By the use of strong magnetic gradient fields (what are known as diffusion gradients), that, in this technique, are applied permanently or also in pulses in addition to the spatially coded gradient fields described above, a collective diffusion movement of the respective molecules (in particular water) becomes noticeable as a weakening of the magnetic resonance signal. Regions in which diffusion ensues are more or less indicated as dark regions in actual MRT images, depending on the strength of the diffusion, which is dependent on various factors. The precise theory of the signal origin in diffusion-weighted imaging is explained below.
A problem in the early development of diffusion-weighted imaging was the marked sensitivity to non-diffusion-associated movements, such as heart motion, respiratory motion and so forth, and the motions connected with these, such as for example brain pulsation (movement of the brain in cerebrospinal fluid). The use of diffusion imaging as a clinical examination method has for the first time made possible the continuous development of faster measurement techniques, such as for example echoplanar imaging (EPI). EPI is a markedly faster measurement method in MRT. Given the use of single-shot echo-planar imaging (SSEPI) sequences, image artifacts that ensue due to unpreventable movement types can be reduced or prevented. Movements as ensue in conventional diffusion-weighted imaging sequences can be effectively “frozen” with SSEPI. A disadvantage caused by the type of phase coding of an SSEPI sequence, however, is the very strong T2*(T2* is the decay duration of the transverse magnetization due to local magnetic field inhomogeneities.) and the very strong phase sensitivity. Both result in strong image artifacts or distortion artifacts, in particular in body imaging with typically short T2 times of human tissue.
“Non-EPI sequences” (generally called steady-state sequences), such as, for example, FISP (Fast Imaging With Steady Precession) and PSIF (the reverse of FISP), generally use the typical spin warp phase coding technique (discrete phase incrementing by means of a phase coding gradient) and are non-sensitive with regard to the artifacts described above. Typically, in such a sequence a monopolar (positive or negative) diffusion gradient (as a rule pulse-type) is switched with a radio-frequency excitation pulse of α<90°. Independent of diffusion gradients, such a radio-frequency pulse α can have three different influences on a magnetization vector:                1. tilting the magnetization vector by the flip angle α relative to the longitudinal direction (z-axis),        2. inverting the magnetization vector by 180°, and        3. no effect at all on the magnetization vector.        
As is explained below in detail, due to these three properties of the radio-frequency excitation pulse, different branched phase curves of the length magnetization and transverse magnetization (also called “echo paths”) occur, each of which exhibits a different diffusion time Δi. Dependent on the phase path or the phase history (which again experiences an expansion via an applied diffusion gradient) the corresponding T1-weighting and T2 weighting is also a specific echo path during the entire phase curve. Overall, an MRT signal results which is formed by a combination of a number of echo paths and therefore exhibits a complexity that can no longer be calculated.
A diffusion-weighted MRT image results from the DADC value (Apparent Diffusion Coefficient, ADC coefficient) (determined per pixel) characterizing the diffusion, which can be calculated from the measurement signals of the employed sequence as well as from the b-value characterizing the measurement. For this, in a diffusion-weighted PSIF sequence with monopolar diffusion gradient pulses in the integration of all echo paths, the respective T1 values and T2 values, as well as the flip angle distribution α(z) of the real radio-frequency pulse used over the selected slice, must be known exactly (R. Buxton, J. of Magnetic Resonance in Medicine 29, 235-243 (1993)). This is not the case in the previously used (steady-state) sequences, thus a precise calculation of the DADC value is not possible in the context of an exact T2 weighting. The results are diffusion-weighted images that exhibit severe artifacts.
For this reason, Y. Zur, E. Bosak, N. Kaplan propose in the Journal of Magnetic Resonance in Medicine 37, 716-722 (1997) to use a bipolar diffusion gradient in place of a monopolar diffusion gradient. The expansion effected by the diffusion gradients thereby is compensated by the phase curves. The diffusion time Δ is thereby well defined, and an ADC calculation from two diffusion measurements is approximately possible with diffusion gradients differing with regard to amplitude.
Admittedly, it is known that, given use of bipolar diffusion gradients, the calculated diffusion coefficients are still strongly dependent on the respective T2 value, which likewise leads to significant artifacts (S. Ding, H. Trillaud et al., J. of Magnetic Resonance in Medicine 34, 586-595 (1995)).
M. H. Cho and C. H. Cho have theoretically shown in the conference volume “Society of Magnetic Resonance in Medicine”, p. 911, Amsterdam (1989) that this T2 weighting can be eliminated given use of a bipolar diffusion coefficient, when a corresponding FISP signal (S+) is known and a corresponding formula is specified for the measured PSIF signal (S−).